The present invention relates to the radio frequency signal handling arts. It finds particular application in conjunction with medical magnetic resonance imaging systems and will be described with particular reference thereto. It is to be appreciated, however, that the invention will also find application in conjunction with other types of magnetic resonance imaging systems, magnetic resonance spectroscopy systems, and the like.
In magnetic resonance imaging, a strong uniform static magnetic field B.sub.0 is generated, often by a superconducting magnet. The static magnetic field B.sub.0 polarizes the nuclear magnetic spin system of an object to be imaged. Superconducting magnets are commonly wound on a cylindrical body former mounted in an annular helium vessel surrounded by an annular vacuum vessel for thermal isolation. The superconducting magnet generates the static magnetic field, B.sub.0, along its own longitudinal axis and the common longitudinal axis of the cylindrical bore of the vacuum vessel, commonly denoted as the z-axis. Alternately, the B.sub.0 field is generated in an open region between a pair of poles. Often, a ferrous flux return path is provided between the poles remote from the open imaging region.
To generate a magnetic resonance signal, the polarized spin system is first excited by applying a magnetic resonance excitation signal or radio frequency field B.sub.1, perpendicular to the z-axis. This RF field B.sub.1 is typically produced by an RF coil located inside the bore of a bore-type magnet or adjacent the pole of an open magnet and closely conforming thereto to maximize the space available to receive a patient. The RF magnetic field is turned on and off to create short RF pulses to excite and manipulate magnetization in the polarized object in the bore. More specifically, the RF excitation pulses tip the magnetization out of alignment with the z-axis and cause its macroscopic magnetic moment vector to precess around the z-axis. The precessing magnetic moment, in turn, generates a radio frequency magnetic resonance signal that is received by the RF coil in a reception mode. Additional RF pulses are commonly applied to manipulate the resonance to form enhanced signal strength RF echoes which are received by the same RF coil or a local RF coil positioned near a region of interest.
In magnetic resonance imaging, it is advantageous for the RF coil to have high sensitivity, high RF power efficiency, and a high signal-to-noise ratio. Also, the B.sub.1 magnetic field which is generated should be uniform. The sensitivity of the RF coil is defined as the magnetic field B.sub.1 created by a unit current. The signal-to-noise ratio is proportional to the sensitivity and to the square root of the coil quality factor, Q.
To encode a sample spatially, magnetic field gradients are applied after the RF excitation. The gradient magnetic fields are typically applied in pulses to generate magnetic field gradients G.sub.x, G.sub.y and G.sub.z linearly along the x, y, and z-directions, respectively, or other selected coordinate system. The gradient pulses typically are generated by gradient magnetic field coils which are also located inside the bore of a bore-type magnet or adjacent the poles of an open-type magnet. Commonly, the gradient field coils are mounted in back of the RF coil in the bore or on the pole piece.
A radio frequency coil is also used to receive magnetic resonance signals emanating from a patient's body. These receive coils may be local coils, such as for receiving RF signals from a patient's head or the larger whole body RF coils located in the bore or on the pole pieces. The local coils can also be used in a transmit mode. The receive coils are typically operable in a quadrature mode. Some local coils include an array or other plurality of RF coils. The signals are then demodulated by a receiver, preferably a digital receiver.
A sequence control circuit controls the gradient pulses and the transmitter to generate a plurality of imaging sequences. For the selected sequence, the receive coil receives one or a plurality of data lines in rapid succession following each RF excitation pulse. Gradient pulses are typically applied before and during reception of resonance echoes and between echoes in multi-echo sequences. An analog-to-digital converter converts each data line to a digital format. Ultimately, the radio frequency signals are demodulated and reconstructed into an image representation by a reconstruction processor which applies a two-dimensional Fourier transform or other appropriate reconstruction algorithm. The image may represent a planar slice through the patient, an array of parallel planar slices, a three-dimensional volume, or the like.
Radio frequency coils in magnetic resonance systems are generally connected to the magnetic resonance system, and more particularly to the RF transmitter and/or the RF receiver of the magnetic resonance system using coaxial cable. Coaxial cable protects the system from picking-up extraneous RF signals which are present in the environment. As is well known, coaxial cable features a surrounding shield or ground conductor separated from a current carrying central conductor by a dielectric material. The surrounding conductor acts as a shield that minimizes the pick-up of foreign frequencies by the cable.
Although coaxial cable is used, there are still coupling problems at resonance frequencies, such as 64 MHZ for hydrogen dipoles in a 1.5 T B.sub.0 field. Among other things, the shield conductor of the coaxial cable itself tends to carry foreign induced currents, such as from TV transmissions, stray harmonics from the gradient pulse oscillators and clocking circuits in nearby equipment, and the like. The induced current is often referred to as "skin current" because it flows on the outside of the shield conductor. The stray RF current tends to flow out of the bore and into other circuits, such as the amplifiers, analog-to-digital converters, receiver, and reconstruction processor to contribute errors in the resultant image.
Balance/unbalance ("Balun") circuitry is used as one means for reducing the noise and/or stray RF currents generated due to induced currents in the coaxial cable. Baluns of the prior art consisted of a cable or tube (often on the order of 1.0 m) that fed into and out of a copper shielded box. The balun was tuned to the frequency of interest, such as by a tuning capacitor. The baluns of the prior art are problematic for a number of reasons. First, the baluns are expensive due to the use of special non-magnetic tuning capacitors. Second, these baluns are inaccessible and sealed which prevents tuning for different coil arrangements or frequencies. Third, these baluns are space consuming. In magnetic resonance scanners, there are severe space limitations. In bore type magnets, there is pressure to reduce the magnet diameter for lower cost competing with pressure to enlarge the patient receiving bore. Similarly, in open magnets, there are competing pressures to move pole pieces closer and to enlarge the patient gap. This compresses the space available for RF coils, gradient coils, shims, baluns, and other associated structures.
Compounding the aforementioned disadvantages, magnetic resonance scanners have multiple RF output channels, such as a channel for each quadrature mode, channels for individual coils of an array, or the like. One or more baluns are incorporated in each channel. The multiple connection of parallel baluns reduces their effectiveness to block RF current. Also, space consuming problems are magnified when multiple baluns are used. Finally, multiple baluns multiply the cost.
The present invention contemplates a new and improved balun design which overcomes the above-referenced problems and others.